Systems and methods for photoacoustic opthalmoscopy

ABSTRACT

Various embodiments of the present invention include systems and methods for multimodal functional imaging based upon photoacoustic and laser optical scanning microscopy. In particular, at least one embodiment of the present invention utilizes a contact lens in combination with an ultrasound transducer for purposes of acquiring photoacoustic microscopy data. Traditionally divergent imaging modalities such as confocal scanning laser opthalmoscopy and photoacoustic microscopy are combined within a single laser system. Functional imaging of biological samples can be utilized for various medical and biological purposes.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application claims priority benefit of U.S. Provisional ApplicationNo. 61/160,907, filed Mar. 17, 2009, and to U.S. Provisional ApplicationNo. 61/335,684, filed Jan. 11, 2010, each of which is incorporatedherein by reference in their entirety.

FIELD OF THE INVENTION

Various embodiments of the present invention relate to multimodalmedical imaging. In particular, various embodiments of the presentinvention relate to photoacoustic and laser scanning microscopy.

BACKGROUND OF THE INVENTION

Visual loss is often considered the most feared complication of humandisease, other than death. Among all the causes that lead toirreversible vision loss, diabetic retinopathy remains a leading cause.Diabetic retinopathy is a vascular disorder that occurs as acomplication of diabetes mellitus. It is estimated that more than 10million adults over the age of 40 in the U.S. have diabetes mellitus.Early detection and treatment of diabetes mellitus could lead tosignificant societal healthcare savings and prevent the loss of sightfor millions of people in the U.S. alone.

Currently, clinical treatment for late stage diabetic retinopathy isassociated with unavoidable side effects that include diminishedperipheral and night vision as well as loss of vision. As diabeticretinopathy progresses the blood vessels supplying oxygen to the eyebecome blocked, which leads to decreased oxygen to the retina. Treatmentof diabetic retinopathy has increasingly been placed upon detectingprogression of the disease at an earlier stage.

SUMMARY

Briefly, in one aspect the invention provides an ophthalmic imagingsystem including a laser capable of generating a laser beam forirradiating a biological sample, an optical coherence tomography (OCT)apparatus capable of generating an OCT probing light beam and adual-axis scanner for scanning both the laser beam and OCT probing lightbeam. Additionally, the invention includes an optical apparatus fordelivering the laser beam and OCT probing light beam to the biologicalsample and an ultrasonic transducer for measuring laser inducedultrasonic waves in the biological sample.

According to another embodiment, the invention provides a method ofimaging a biological sample, which includes generating a laser beamcapable of irradiating a region of interest on a biological sample, thelaser beam generated from a tunable laser system, merging the laser beamwith an optical coherence tomographic probing light to form a mergedbeam and scanning the merged beam using a dual-axis optical scannercapable of two-dimensional raster scanning. The method also includesrecording photoacoustic waves generated by the irradiated retinalregion, recording light reflected off of the biological sample,controlling the recording and scanning within a single time base andgenerating an ophthalmic image based at least in part upon scanning themerged beam and recording photoacoustic waves and reflected light.

In another aspect, the invention includes an OCT-guided laser scanningphotoacoustic microscope with a tunable laser capable of irradiating abiological sample with a laser beam, a dual-axis galvanometer for rasterscanning the laser beam and a fiber-based spectral-domain opticalcoherence tomography (OCT) system capable of generating an OCT probinglight beam and the OCT system having a spectrometer for detectinginterference signals within a spectral domain. The invention alsoincludes an optical delivery system for merging the laser beam and OCTprobing light beam and delivering a merged beam to the biologicalsample. Additionally included, is an ultrasound transducer integratedwithin a contact lens for detecting photoacoustic waves generated by theirradiated biological sample and a GUI for displaying a image of thebiological sample based at least in part upon the photoacoustic wavesand spectrometer detected signals.

Other aspects of the invention will become apparent by consideration ofthe detailed description and accompanying drawings.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a system diagram of an exemplary photoacoustic opthalmoscopysystem in accordance with at least one embodiment of the presentinvention;

FIG. 2 is a system diagram of another exemplary photoacousticopthalmoscopy system in accordance with at least one embodiment of thepresent invention;

FIG. 3 is a flow chart representing a method for noninvasive ophthalmicimaging in accordance with at least one embodiment of the presentinvention;

FIG. 4 is a system diagram of an exemplary confocal photoacousticopthalmoscopy system in accordance with at least one embodiment of thepresent invention;

FIG. 5 is a flow chart representing another method for noninvasiveophthalmic imaging in accordance with at least one embodiment of thepresent invention;

FIG. 6 is an OCT guided photoacoustic opthalmoscopy system in accordancewith at least one embodiment of the present invention;

FIG. 7 is a functional flow diagram of an optical delivery and scanningsystem in accordance with at least one embodiment of the presentinvention;

FIG. 8 is another OCT guided photoacoustic opthalmoscopic system inaccordance with at least one embodiment of the present invention;

FIG. 9 a is a side perspective view of a contact lens in combinationwith a needle ultrasound transducer in accordance with at least oneembodiment of the present invention;

FIG. 9 b is a side perspective view of the contact lens according toFIG. 9 a, the contact lens having a hollow center;

FIG. 10 a is a side perspective view of a contact lens in combinationwith a ring transducer in accordance with at least one embodiment of thepresent invention; and

FIG. 10 b is a top plan view of the contact lens as shown in FIG. 10 a.

DETAILED DESCRIPTION

Before any embodiments of the invention are explained in detail, it isto be understood that the invention is not limited in its application tothe details of construction and the arrangement of components set forthin the following description or illustrated in the following drawings.The invention is capable of other embodiments and of being practiced orof being carried out in various ways. Also, it is to be understood thatthe phraseology and terminology used herein is for the purpose ofdescription and should not be regarded as limited. The use of“including,” “comprising” or “having” and variations thereof herein ismeant to encompass the items listed thereafter and equivalents thereofas well as additional items. The terms “mounted,” “connected” and“coupled” are used broadly and encompass both direct and indirectmounting, connecting and coupling. Further, “connected” and “coupled”are not restricted to physical or mechanical connections or couplings,and can include electrical connections or couplings, whether direct orindirect. Also, electronic communications and notifications may beperformed using any known means including direct connections, wirelessconnections, etc.

It should be noted that a plurality of hardware and software baseddevices, as well as a plurality of different structural components maybe utilized to implement the invention. Furthermore, and as described insubsequent paragraphs, the specific configurations illustrated in thedrawings are intended to exemplify embodiments of the invention and thatother alternative configurations are possible. Further, use of the term“processor” is meant to include systems architected in such a way as tohave only a single processor or multiple distributed processorsfunctioning serially or in parallel to perform the processing functions.In addition, in multi-processor systems, the description is meant toencompass constructions having all processors in one machine or in onelocation, or constructions having processors distributed acrossdifferent machines or locations.

FIG. 1 shows a laser-scanning optical resolution photoacoustic imagingsystem 100 is provided. The system 100 includes a tunable pulse laser102, an optical delivery and scanning apparatus 104, a processor 106, anultrasonic detector 108, an amplifier 110 and graphical user interface(GUI) 112.

The system 100 is a multimodal imaging system that combinesphotoacoustic microscopy with laser scanning and optical resolutionmicroscopy. A laser beam 114 is delivered to a biological sample 116 bythe optical delivery and scanning apparatus 104. The laser beam 114 is asource for irradiating the biological sample 116. Absorbed photons ofthe irradiated biological sample 116 are detected by the system 100within high spatial resolution. The system 100 performs a hybridphotoacoustic imaging method that detects the laser-induced ultrasonicwaves to acquire distributions of internal optical energy depositions.When laser pulses irradiate biological tissues, optical energy isabsorbed by, for example, blood vessels and converted to heat. In manycases, there is a milli-degree temperature rise. The temperature rise isfollowed by thermoelastic expansions in the tissue of the biologicalsample 116. These expansions create wideband ultrasonic waves, which areoften referred to as photoacoustic waves. The photoacoustic waves aredetected and used to quantify the optical absorption properties of thebiological sample. Additionally, the photoacoustic waves are used toform an image of the biological sample 116 based upon the opticalabsorption contrast of the tissue. The resulting image is displayed onthe GUI 112.

The biological sample 116 in the present embodiment is a human eye.However, it is contemplated that functional images can be obtained of amultitude of different biological samples 116. By example, thebiological sample can include cells and molecules in suspension,physiological appendages, small animal organs, including ears, skin,eyes, brain, and internal organs, and human eyes and skin. In thepresent embodiment, the laser beam 114 enters through the pupil and isdirected to a retinal region of interest, which is identified as thefield of view (FOV) 118, within the eye 116. The FOV is also the regionbeing targeted by the ultrasonic transducer 108.

The ultrasonic transducer 108 is kept stationary while information isacquired. Data acquired by the ultrasonic transducer 108 is delivered tothe amplifier 110 and then to the processor 106. The processor 106performs a plurality of functions, and it is contemplated that more thanone processor 106 can be employed within the system 100. The processor106 includes executable code that is capable of performing a variety offunctions. In particular, the processor is capable synchronizing thelaser triggering and operation of the optical scanner 132. Some of thevarious functions performed by the processor 106 include controllingwhen the laser 102 is triggered and controlling the delivery andscanning apparatus 104, which directly effects how the laser beam 114 isdelivered. The processor 106 can be connected to a signal digitizer thatstores detected digital signals for image formation.

The processor 116 can be more than one computing device, or a singlecomputing device having more than one microprocessor. It is contemplatedthat the processor 106 is a stand alone computing system with internalor external memory, a microprocessor and additional standard computingfeatures. The processor 116 can be selected from the group comprising aPC, laptop computer, microprocessor, or alternative computing apparatusor system. The processor 116 receives imaging data that has beenobtained through photoacoustic microscopy and laser scanning opticalresolution microscopy and generates an image of a biological samplebased upon these and other data inputs.

FIG. 2 illustrates another embodiment of the laser scanning opticalresolution photoacoustic imaging system 100. Like parts are identifiedusing like reference numerals The tunable pulse laser system 102 (See.FIG. 1) includes a neodymium-doped yttrium lithium fluoride (Nd:YLF)laser apparatus 120, a dye laser 122 and a fast photodiode 124. Anexemplary laser apparatus 120 includes a model IS8II-E produced by EdgeWave GmbH, Germany and an exemplary dye laser 122 includes a Cobra modelproduced by Sirah Laser and Plasmatchnik GmbH, Germany.

The dye laser 122 is pumped by the laser apparatus 120 and functions asthe irradiation source. The pulse duration of the laser apparatus 120 isabout 6 ns. However, it is contemplated that the pulse duration canrange from about 1 ns to about 10 ns. Additionally, it is contemplatedthat the pulse duration can be less than 1 ns and greater than 10 ns.The laser 120 is capable of operating with a pulse repetition rate in arange of about 3 kHz to about 6 kHz. Alternatively, the pulse repetitionis less than about 3 kHz and greater than about 6 kHz. The laser 120 ispreferably within an optical tuning range of about 500 nm to about 700nm. Alternatively, the optical tuning range is from about 300 nm toabout 1300 nm. It is further contemplated that the tunable pulse lasersystem 102 is replaced with a plurality of individual fixed wavelengthlasers, the wavelengths being within the ranges identified above, thatare controlled by the processor 116.

The photodiode 124 relays laser beam 114 information to the processor106, such as detecting the laser pulses. In addition, the photodiodeprovides information to the processor 106, which is capable oftriggering the acquisition of data to avoid the impact of laserjittering. The energy of each laser pulse is recorded by the photodiode124 and sent to the processor 106 for storage in a memory storage device(not shown). Compensation for pulse energy instability can be achievedby utilizing energy of each laser pulse. An exemplary fast photodiodeincludes a model DET10A manufactured by Thorlabs, Newton, N.J.

The optical delivery and scanning system 104 includes an iris 126, abeam expander 128, an attenuator 130, a dual-axis scanner 132 and anobjective lens 134. The laser beam 114 is spatially filtered by the iris126 and expanded by the beam expander 128. (Model BE03M-A, Thorlabs) Inthe present embodiment, the attenuator 130 is a neutral density filter.(Model FW2AND, Thorlabs) The expanded and attenuated laser beam 114passes through the dual-axis scanner 132, which is, by example, an x-ygalvanometer (Model 6230H, produced by Cambridge Technology Inc.,Lexington, Mass. and Model QS-10, Nutfield Technology). After passingthrough an objective lens 134 the laser beam is focused on a biologicalsample 116 to be irradiated. Alternatively, the optical scanner 132 canbe a polygon mirror scanner or a combination of a dual-axis scanner anda polygon scanner.

In another embodiment, the optical delivery system 104 is based ontraditional optics. With traditional optics the laser light beam can beexpanded and collimated to cover a biological sample, such as the pupilof an eye, whether the eye is dilated or undilated. The cornea and lensof the eye focus the collimated light on to a retinal region ofinterest. To compensate for the refractive error of the eye, for examplemyopia, the position of the objective lens of the optical deliverysystem 104 is adjustable to ensure adequate focus of the light onto theretinal region of interest. Considering the size of an average humanpupil (about 7 mm in diameter) and the aberration of the eye, theexpected optical focusing spot is about 10 μm. The lateral resolution ofthe present embodiment is about 10 μm. Alternatively, the lateralresolution can be adjusted less than about 10 μm and greater than about10 μm.

In yet another embodiment, the optical delivery system 104 is based uponadaptive optics that includes a wavefront sensor, a deformable mirrorand processor (not shown). The light source is either the existing pulselaser beam 114 or an additional laser light coupled to the system 104.Light reflected from the fondus of an eye is partially reflected by thebeam splitter into the wavefront sensor, which detects the disturbanceof the wavefront sensor caused by the aberrations of the eye and opticaldelivery system. Based at least in part upon the output of the wavefrontsensor the processor computes the compensation and transfers thecompensation signal to the deformable mirror. The deformable mirrorapplies phase modulations to the laser beam to compensate for thedisturbance.

Two-dimensional scanning of the optical focus is achieved by a dual-axisgalvanometer, which can be performed regardless of the various opticaldelivery methods discussed herein. The dual-axis optical scanner 132 iscapable of various types of scanning, such as two dimensional scanning,raster scanning, spiral scanning and circular scanning. According to atleast one embodiment of the present invention, the ultrasonic detectorsare kept stationary on the eyelids of a human subject whiletwo-dimensional optical scanning is performed. Ultrasound gel is usedfor better coupling between the detector and the eyelid. Since theultrasound detectors 108 are kept stationary, the imaging FOV is theultrasound detection region. The center frequency of the ultrasonictransducer is chosen according to the diameters of the major retinalvessels, which range from about 20 μm to about 100 μm. The frequencyranges from about 10 MHz to about 50 MHz. The bandwidth of the detector108 is preferably greater than about 60% in order to provide high axialresolution.

The ultrasound detector 108 detects and measures laser inducedultrasonic waves generated by an irradiated biological tissue. The datadetected and measured by the ultrasound transducer 108 is transmitted tothe processor 106 for ultimately generating an image of the biologicalsample based at least in part upon the photoacoustic waves, anddisplayed on the GUI 112. In the present embodiment, the ultrasoundtransducer 108 has a stationary position with respect to the biologicalsample 116 and is selected from the group comprising a single stationarydetector, an array of detectors, a contact lens integrated with anultrasound detector and a contact lens integrated with an array ofdetectors. The transducer 108 is selected to operate in a range of about10 MHz to about 50 MHz. Alternatively, the transducer operates at lessthen about 10 MHz or greater than about 50 MHz.

FIG. 3 illustrates an exemplary method of noninvasive ophthalmicimaging. The system 100 (See FIG. 2.) is initiated at step 136 and alaser beam is generated from a tunable laser system at step 138. Thelaser beam is capable of irradiating a biological sample, which in turngenerates photoacoustic waves. The laser beam 114 transmitted from a dyelaser is spatially filtered by an iris at step 140. This is followed byscanning the laser beam 114 with a dual-axis scanner at step 142, whichprovides two-dimensional scanning of the laser beam within the field ofview for the region of interest on the biological sample. Prior toreaching the biological sample, the laser beam is collimated and focusedon the biological sample with an optical apparatus at step 144. Thebiological sample is irradiated after the focused laser beam isdelivered at step 146. In response to the irradiated biological sample,photoacoustic waves are detected at step 148 and processed at step 150.The photoacoustic waves are amplified by a wideband amplifier at step152, followed by digitizing the photoacoustic signals and storing themwithin the processor 106. Alternatively, the signals are stored by adata acquisition board operatively connected to the processor 106. Basedupon the scanned data and photoacoustic signals the processor generatesa functional image of the biological sample and displays the image onthe GUI 112 at step 154.

In order to measure functional information such as sO₂, photoacoustic(PA) imaging performs multi-wavelength measurements within properoptical spectral ranges. This is in the same way as does NIRS, where HbRand HbO₂ are treated as the dominant optical absorbers at eachwavelength (λ_(i)). Thus the blood absorption coefficientμ_(a)(λ_(i))(cm⁻¹) can be expressed through Equation Set 1.

μ_(a)(λ_(i))=ε_(HbR)(λ_(i))[HbR]+ε _(HbO) ₂ (λ_(i))[HbO ₂],  EquationSet 1

ε_(HbR)(λ_(i)) and ε_(HbO2)(λ_(i)) are the known molar extinctioncoefficients (cm⁻¹ M⁻¹) of HbR and HbO₂ at wavelength λ_(i),respectively; and [HbR] and [HbO₂] are the concentrations of the twoforms of hemoglobin, respectively. Since the amplitude of the acquiredlocalized PA signal φ(λ_(i),x,y,z) is proportional to the local opticalenergy deposition, the μ_(a)(λ_(i)) can be replaced by φ(λ_(i),x,y,z) tocalculate the [HbR] and [HbO₂] in relative values. Least-squares fittingleads to Equation Set 2:

$\begin{matrix}{{{\begin{bmatrix}\lbrack{HbR}\rbrack \\\left\lbrack {HbO}_{2} \right\rbrack\end{bmatrix}_{({x,y,z})} = {\left( {M^{T}M} \right)^{- 1}M^{T}{\Phi \left( {x,y,z} \right)}K}},{Where}}{{M = \begin{bmatrix}{ɛ_{HbR}\left( \lambda_{1} \right)} & {ɛ_{{HbO}_{2}}\left( \lambda_{1} \right)} \\\vdots & \vdots \\{ɛ_{HbR}\left( \lambda_{n} \right)} & {ɛ_{{HbO}_{2}}\left( \lambda_{n} \right)}\end{bmatrix}},{{\Phi \left( {x,y,z} \right)} = {\begin{bmatrix}{\varphi \left( {\lambda_{1},x,y,z} \right)} \\\vdots \\{\varphi \left( {\lambda_{n},x,y,z} \right)}\end{bmatrix}.}}}} & {{Equation}\mspace{14mu} {Set}\mspace{14mu} 2}\end{matrix}$

K is the proportionality coefficient that is related to the ultrasonicparameters and the wavelength-dependent change of the local opticalfluence as light passes through the skin. Thus the sO₂ image iscalculated using Equation Set 3.

$\begin{matrix}{{sO}_{2_{({x,y,z})}} = \frac{\left\lbrack {HbO}_{2} \right\rbrack_{({x,y,z})}}{\left\lbrack {HbO}_{2} \right\rbrack_{({x,y,z})} + \lbrack{HbR}\rbrack_{({x,y,z})}}} & {{Equation}\mspace{14mu} {Set}\mspace{14mu} 3}\end{matrix}$

Due to the unknown coefficient K, only relative concentration of the HbRand HbO₂ are calculated from Equation Set 2. However, the sO₂ fromEquation Set 3 is an absolute measurement. Although two wavelengths areenough to determine SO₂ in principle, it is recommended to use morewavelengths in order to reduce the influence of measurement error.Published molar extinction coefficients of HbR and HbO₂ are used inEquation Set 2.

Various embodiments of the present invention provide a laser scanningoptical resolution photoacoustic microscope 100. Higher scanning speedis enabled by the laser scanning systems described above. Higherscanning translates into increased data acquisition over previouslyknown imaging modalities. Additionally, complete raster scanning andother complex scanning patterns such as concentric circular scanning andtwo dimensional arc scanning are available with the present system 100that allow localized measurements of single vessels that translates intoimproved visualization and detection of maladies within the vessels.Experiments relating to various embodiments of the present invention arediscussed within the article Laser-Scanning Optical ResolutionPhotoacoustic Microscopy, Optics Letters, Vol. 34, No. 12, herebyincorporated by reference in its entirety herein.

FIG. 4 illustrates a confocal scanning laser photoacoustic microscopesystem 156. The system 156 includes similar elements as the system 100provided in FIG. 3, but utilizes a confocal scanning laser microscopicapparatus 158. The apparatus 158 includes an optical coupler 160, asecond photodiode 162 and a third photodiode 164. The optical coupler160 includes a first output arm 166 and a second output arm 168.Although a single processor is contemplated, the present embodimentutilizes a scan control processor 170 and a data acquisition processor172.

The exemplary system 156 generally provided in FIG. 4 can include avariety of exemplary elements. By example, a tunable dye laser 122pumped by a ND:YLF laser apparatus 120 having a pulse duration of about6 ns can be used as a biological sample 116 irradiation source. Theoutput wavelength of the irradiation source can vary depending uponsystem 156 parameters. By example, the wavelength is selected in a rangeof about 500 nm to about 900 nm. It is contemplated that the wavelengthcan be less than 500 nm and greater than 700 nm.

The laser beam 114 is transmitted by the dye laser 122 and spatiallyfiltered by an iris 126 and attenuated by a neutral density filter (notshown) before entering a confocal scanning laser microscopic apparatus158. The apparatus 158 includes a 2×2 single-mode optical fiber coupler160 (Model FC-632, Thorlabs Inc., Newton, N.J.) operatively connected tothe laser 122, a second photodiode 162 and a third photodiode 164. Theoutput light from the optical coupler first arm 166 is collimated andexpanded. An exemplary diameter is between about 5 mm and 15 mm. Theexpanded laser beam is scanned by a dual-axis galvanometer 132. Anobjective lens 134 is used to focus the laser beam within the region ofinterest 118 upon the biological sample 116. An exemplary objective lensis an achromatic lens with a focal length of about 40 mm. The secondoutput arm 168 directs output light from the optical coupler 158 to thesecond photodiode 162 (Model DET10A, Thorlabs Inc., Newton, N.J.). Thesecond photodiode records the energy of every laser pulse generated bythe laser 122. This energy data is fed to the processor 172 and used tocompensate for energy instability. The detected photoacoustic amplitudeis normalized, by example it is divided, by the detected laser pulseenergy for the compensation. Furthermore, the amplitude of thephotoacoustic signal is proportional to the laser energy.

A confocal image of the biological sample is generated based at least inpart upon the laser light photons reflected from the biological sample.The photons are detected by the third photodiode 164 and transmitted tothe processor 172. An exemplary third photodiode 164 is a 10 MHz Siphotodiode (Model 2107-FC, New Focus Inc., Santa Clara, Calif.).Generation of confocal images from a fiber optic coupler can begenerated by several processes previously known. By example, theprocessor 106 undertakes various calculations and processes to generatean image, such as plotting the detected optical intensity against thespatial location of the optical focus, thereby generating an image.Acquisition of data for the biological sample was triggered by laserpulses directed by the first photodiode 124, which is designed to avoidlaser jittering.

In parallel with the confocal image generation an ultrasonic transducer108 detects induced photoacoustic waves emanating from the biologicalsample. The photoacoustic waves are amplified by the amplifier 110 andtransmitted to the processor 172. The processor 172 generatesphotoacoustic microscopic images based at least in part upon thephotoacoustic waves. The processor 172 merges the photoacoustic andconfocal based images to provide a single multimodal functional image ofa biological sample.

The exemplary system 156 and the elements identified above represent thecombination of a fiber optic confocal microscope and a laser scanningoptical resolution photoacoustic microscope. The system 156 isconfigured to generate a combined ophthalmic image based at least inpart upon a scanned ophthalmic region of interest, photons reflectedfrom the ophthalmic region of interest and photoacoustic waves generatedfrom an irradiated ophthalmic region of interest. As shown in FIG. 4,the system 156 uses a single laser light source 122. In anotherembodiment, it is contemplated that multiple laser light sources can beintegrated within the system 156.

The system 156 configuration was designed to permit multimodalophthalmic imaging, specifically the novel integration of laser scanningoptical resolution microscopy and a fiber optic confocal microscopy. Inparticular, these traditionally divergent technologies were integratedwith the use of a single laser light source, thereby allowingsimultaneous imaging of a biological sample based at least in part uponoptical absorption and scattering contrasts.

The exemplary system 156 is also referred to as a multimodal ophthalmicimaging system. As described above, the system 156 provides aphotoacoustic microscope elements for generating functional images of abiological sample. The optical coupler 158, photodiodes 162,164 andprocessor 172 combine to form a fiber optic confocal microscope forgenerating functional images of the biological sample. The processor 172registers the photoacoustic and confocal generated images based at leastin part upon a single laser source to provide an enhanced functionalimage of the biological sample, which can be displayed on a GUI 112.

FIG. 5 illustrates an exemplary method of noninvasive ophthalmic imagingusing a confocal scanning laser microscope. The system 156 (See FIG. 4)is initiated at step 174 and a laser beam is generated from a tunablelaser system at step 176. The laser beam is capable of irradiating abiological sample, which in turn generates photoacoustic waves. Thelaser beam transmitted from a dye laser is spatially filtered by an irisat step 178, followed by scanning the laser beam with a dual-axisoptical scanner at step 180, which causes two dimensional scanning ofthe laser beam within the field of view for the region of interest onthe biological sample. Additional scanning techniques are contemplated,such as raster scanning Prior to reaching the biological sample, thelaser beam is collimated and focused on the biological sample with anoptical apparatus at step 182. The biological sample is irradiated afterthe focused laser beam is delivered at step 184. Photons reflected fromthe biological sample are collected at step 186 and processed at step188. In response to the irradiated biological sample, photoacousticwaves are detected at step 190 and processed at step 192. Thephotoacoustic waves are amplified by a wideband amplifier at step 194,followed by digitizing the photoacoustic signals and storing them withinthe processor 106. Alternatively, the signals are stored by a dataacquisition board operatively connected to the processor 106. Based atleast upon the processed photons and photoacoustic signals, theprocessor 172 generates a functional image of the biological sample anddisplays the image on the GUI 112 at step 196.

The biological sample can be selected from a plurality of differentanimals and organs. In particular, an exemplary biological sample is aneye to be noninvasively imaged by the present embodiment. In anotherembodiment, mice ears have been imaged using the system 156 describedabove and reported within the following articles: IntegratedPhotoacoustic Microscopy and Fiber-Optic Confocal Microscopy Using aSingle Laser Source, Proceedings of OSA BIOMED, Miami, Fla. (2010) andCollecting Back-Reflected Photons in Photoacoustic Microscopy, OpticsExpress 18, 1278-1282 (2010), both of which are hereby incorporated byreference in their entirety herein.

System 156 (See FIG. 4) can also be used to for ophthalmic imaging thatincludes irradiating a retinal region of interest with a laser beamgenerated from the tunable laser 122. The retinal region of interest isthen scanned using a dual-axis galvanometer capable of two-dimensionalraster scanning, which is followed by collecting photons reflected fromthe retinal region of interest with a 2×2 single-mode fiber opticalcoupler. The photoacoustic waves generated by the irradiated retinalregion are then collected by an ultrasound transducer. The processor 172controls the collection of the photons, recording of the laser pulsesand scanning the laser beam within a single time base. Based at least inpart upon the scanned retinal region of interest, photons reflected fromthe retinal region of interest and recorded photoacoustic waves afunctional image of the retina is generated. The present system 156 iscapable of combining O2 consumption data, metabolic information andblood flow velocity for purposes of generating functional imaging.

FIG. 6 illustrates an optical coherence tomography (OCT) guided scanninglaser optical resolution photoacoustic microscopic system 198. Thesystem 198 utilizes many of the photoacoustic microscopy elements asprovided within system 100 (See FIG. 1). The present system 198incorporates spectral-domain OCT with laser scanning photoacousticmicroscopy to provide a method of enhanced imaging of a biologicalsample 116.

The OCT-guided laser scanning photoacoustic microscope 198 includes atunable laser 120 capable of irradiating a biological sample 116 with alaser beam, a dual axis galvanometer 132 for raster scanning the laserbeam, and a fiber-based spectral-domain optical coherence tomography(OCT) system capable of generating an OCT probing light beam. The OCTsystem includes a spectrometer 212 for detecting interference signalswithin a spectral domain. Additionally, an optical delivery system formerging the laser beam and OCT probing light beam and delivering amerged beam to the biological sample is provided. The system 198 alsoincludes an ultrasound transducer integrated within a contact lens fordetecting photoacoustic waves generated by the irradiated biologicalsample 116. A GUI 112 for displaying an image of the biological samplebased at least in part upon the photoacoustic waves and spectrometerdetected signals is also provided within the system 198.

The imaging system 198 can be used for ophthalmic imaging by irradiatinga retinal region of interest with a merged laser beam generated from atunable laser and an OCT probing light, which. is scanned together usinga dual-axis optical scanner capable of two-dimensional raster scanning.The system controls the recording of photoacoustic waves and scanning aretinal region of interest within a single time base. Then, anophthalmic image is generated based at least in part upon the recordingof photoacoustic waves and scanning a retinal region of interest.

The system 198 includes a spectral domain OCT system 200 capable ofgenerating an OCT probing light beam. The system 200 includes a superluminescent diode light source 202, a fiber coupler 204, a reference arm206, a sample arm 208 coupled to the photoacoustic apparatus 210 and aspectrometer 212 for detecting interference signals within the spectraldomain. The system 198 also includes an optical scanning apparatus 214for scanning both a laser beam and OCT probing light, as well asdelivering the combined or merged light to the biological sample 116(See FIG. 7). Alternatively, OCT systems presently known and suitablefor biological sample imaging can be used for integration with thephotoacoustic microscope elements of the system 198. Another embodimentis provided in FIG. 8, which discloses a OCT guided laser scanningphotoacoustic microscope system 198.

The spectrometer 212 includes a line scan CCD camera that acquires OCTimages. By example, a CCD camera can be configured with an exposure timeof 36 μs, which can acquire OCT images with a line rate of 24 kHz whilehaving a measured sensitivity greater than about 95 dB. The calibrateddepth resolution is about 6 μm in the biological sample and a lateralresolution of about 20 μm. It is contemplated that alternative OCTtechniques can be used, for example, time-domain OCT and swept-laserOCT. The OCT systems can also be implemented in free space.Additionally, it is contemplated that varying exposure time, tissuedepth resolution and lateral resolutions can be obtained through use ofthe spectrometer 212 and similar alternatives.

An exemplary system 198 includes a frequency-doubled Q-switched ND:YLGlaser 218 (Model SPOT-10-100-532, Elforlight Ltd., UK) having a 532 nmwavelength, 10 μJ/pulse, 2 ns pulse duration and 30 kHz pulse repetitionrate is used as an illumination source. The out put laser 218 light isattenuated with a series of neutral density filters (not shown) beforebeing coupled to a 1×2 single mode optical beam splitter 216. The outputlaser light from a first arm 220 of the splitter 216 is combined with anOCT light beam in the optical scanning apparatus 214. The second arm 222of the splitter 216 is connected to a multimode fiber beam splitter 224.The two outputs of the splitter 224 are connected to a first photodiode226 and a second photodiode 228.

Induced photoacoustic waves of the irradiated biological sample 116 aredetected by the ultrasound transducer 108. The signals are amplified bythe amplifier 110 and transmitted to a processor/computer 106.Processing of the photoacoustic data is performed as discussed withinsystem 100 herein.

The OCT and photoacoustic aspects of the system 198 are synchronized bythe processor 170, 172 despite the difference in the achievable imagingspeed of the respective aspects. Automatic registration of the OCT andphotoacoustic images is obtained by controlling the timing, within asingle time base, of the laser triggering, acquisition of thephotoacoustic data, dual-axis galvanometer scanning and OCT dataacquisition. By example, an analog-output board (Model PD2-AO-16, UnitedElectronic Industries) can be used to trigger data acquisition by thespectrometer 212 (CCD Camera), the laser 218 and dual-axis galvanometer132. Photoacoustic data acquisition can be triggered by the photodiode228 in order to avoid laser jittering.

FIG. 7 illustrates a more detailed view of the optical scanningapparatus 214. The apparatus 214 fuses OCT light beams 230 andphotoacoustic light beams 232. The light beams are combined with a beamcombining dichrotic mirror 234. After combination, the light beams arescanned together with a dual axis scanner 132. The scanner 132 andphotoacoustic laser triggering/data acquisition are synchronized asdescribed above. The apparatus 214 further includes a relay lens 236, amirror 238 and a yolk lens 240.

During image alignment, the laser 218 is turned off and an area ofinterest is selected through guidance of the OCT image. The OCT image isoptimized to ensure good focus of the light upon the retinal region ofinterest in the eye 116. The system is activated for image acquisitionmode once proper alignment of the image is obtained, which includes OCTand photoacoustic based images to be obtained. The same scanning andcontrolling system 214 is used for obtaining images from the differentimage modalities.

More than one mode of operation for the combined imaging system 198 iscontemplated. By example, a first mode includes a strategy referred toas “imaging mode”. The imaging strategy includes the acquisition of bothOCT images and photoacoustic based images for an entire imageacquisition period. This includes three-dimensional registration of OCTand photoacoustic images. Images for a relatively large area or a singleblood vessel are obtained through this imaging mode. This imaging modeallows for two registered three-dimensional images based upon differentcontrast mechanisms to be acquired simultaneously. Spatial hemoglobinoxygen distribution can be acquired together with flow distribution andvessel morphology. As a result, metabolic rate for any specificallyimaged retinal region can be quantified. Photoacoustic images for thepresent system refer to images obtained through use of laser scanningoptical resolution photoacoustic microscopy.

An exemplary second imaging mode is referred to as a “sampling mode”,whereby OCT images are obtained for an entire image acquisition period,while photoacoustic images are obtained only when vessels within thebiological sample are scanned. The sampling mode provides OCTthree-dimensional images and hemoglobin oxygen saturation levels for thevessels based upon the photoacoustic imaging.

The fusion of OCT-based and photoacoustic-based images of the biologicalsample 116 allows both the anatomical and microvasculature to bevisualized in a single volumetric image. The system 198 allows forseamless integration of two images along the lateral directions. Inorder to register two B-scan images along the axial direction the depthrelationship between the photoacoustic and OCT based images must bedetermined. Then, the photoacoustic volumetric image is shifted alongthe axial direction to register with the OCT based volumetric image. Inanother embodiment, mice ears have been imaged using the system 198described above and reported within the following articles: SimultaneousMultimodal Imaging with Integrated Photoacoustic Microscopy and OpticalCoherence Tomography, Optics Letters, Vol. 34, No. 19 (2009) andNaturally Combined Photoacoustic Microscopy and Optical CoherenceTomography for Simultaneous Multimodal Imaging, Optics Letters, 34,2961-2963 (2009), both of which are hereby incorporated by reference intheir entirety. Additionally, the following article discusses ophthalmicimaging using OCT guided laser scanning photoacoustic microscopy:Photoacoustic Ophthammoscopy for in vivo Retinal Imaging, OpticsExpress, 18, 3967-3972 (2010), which is hereby incorporated by referencein its entirety herein. (All articles referenced and incorporated hereinrepresent the applicant's own work.)

In another embodiment, optical absorption and autofluorescence can beobtained through a modified OCT guided laser scanning optical resolutionphotoacoustic microscope. A change from the fiber optic based confocalscanning laser ophthalmascope to a free space is made. This modificationavoids direct reflection from the fiber tip. Directly reflected lightcould potentially be stronger than collected fluorescence light. Aphotomultiplier tube (PMT) or an avalanche photo-diode in combinationwith a dichroic mirror, a band pass filter and a pin-hole are utilized(not shown). The fiber based detection system can be changed tofree-space based. An optical filter and a Dichotic mirror will route thefluorescence photons to a pin-hole and then be detected by ahigh-sensitivity photon detector such as amplified photodiode, photonmultiplier tube, and avalanche photodiode. For auto-fluorescence imagingthe lateral position is provided by the scanning of the stimulatinglight. Auto-fluorescence imaging provides information about thedistribution and concentration of lipofuscin and melanin, two importantpigment components related to the function of RPE.

FIGS. 9-10 show an exemplary device 242 including a contact lens 244combined with an ultrasound transducer 246. The contact lens 244 isconfigured for placement directly on an eye to be imaged and theultrasonic transducer 246 is operatively connected to the contact lens244. The contact lens includes an outer wall 248 to an inner wall 250,the lens having a substantially uniform thickness measured by a distancebetween the outer wall 248 and inner wall 250. FIG. 9 provides twoembodiments of the device 242. In both embodiments the transducer 246 isa needle transducer placed obliquely into the contact lens 244. Thefirst embodiment (FIG. 9 a) includes a standard-shaped contact lens.Preferably the contact lens is powerless and has no refractive power. Inanother embodiment the contact lens is powered and can assist opticalfocusing within the eye. The contact lens can be configured for anindividual's specific ophthalmic physiology. By example, a person withmyopia typically requires corrective lens or contact lens in order tohave 20/20 vision. A contact lens 244 can be powered based upon theparticular myopic structure of an individual's eye. The ultrasoundtransducer 246 is selected from the group comprising a needletransducer, a ring transducer and an ultrasonic array transducer. It iscontemplated that the needle transducer 246 can be positioned inalternative configurations while still obtaining the desiredphotoacoustic data.

The second embodiment (FIG. 9 b) includes the same contact lens as shownin the first embodiment (FIG. 9 a) except there is a cut-out at the apexof the contact lens 244, thereby providing a hollow aperture 252 in thecontact lens 244. For this embodiment, the contact lens is preferablypowerless, as optical illumination reaches the eye directly, unlike theembodiment shown in FIG. 9 a. The hollow center aperture 252 isconfigured and a sized to have a diameter larger than a dilated pupil ofthe eye upon which the device 242 has been placed. The size of thecontact lens 244 is dependent upon the size of the eye, by example, thehollow center 252 has a diameter in a range of about 2 mm to about 8 mmfor use in a human eye. It is contemplated that the size of the contactlens 244 and the hollow center 252 can be selected based upon the sizeof the eye on which the device 242 is placed.

The contact lens 244 can be manufactured from a plurality of suitablematerials that are used for conventional contact lenses, including thosefor refractive correction and fashion. By example, the contact lensmaterial can be selected from the group comprising glass, polymethylmethacrylate, an epoxy composite, materials utilized for hard and softcontact lens, optically transparent and gas permeable materials.

The needle transducer 246 is placed within an orifice that extends froman outer wall 248 to an inner wall 250 of the contact lens 244. Afterplacement of the needle transducer there is a void 252 purposefullyconfigured and located proximal to the inner wall 250. After the device242 is placed directly onto an eye to be imaged the void will fill withfluid from the eye, such as tears. The natural fluid acts as anultrasonic coupling medium for enhanced ultrasonic wave detection.Optical illumination passes through the contact lens before entering theeye. The needle transducer can be selected from a variety ofcommercially available transducers. By example, the needle transducercan have an outer diameter between about 0.5 mm and 1.3 mm, constructedfrom a PMN-PT single crystal and be housed within a polyamide tube.Exemplary transducers generate a center frequency in a range of about 10MHz to about 50 MHz. Alternatively, the center frequency can be lessthan about 10 MHz and greater than about 50 MHz.

It is contemplated that the device 242 has multiple contact lens 244 andtransducer 246 configurations. FIGS. 10 a and 10 b provide two views ofan alternative device 242 configuration. A side perspective view isprovided in FIG. 10 a, and a top plan view of the device 242 is providedin FIG. 10 b. A ring shaped ultrasound transducer is combined with apowerless contact lens with a hollow center spatially configured forcircumferentially positioning with respect to an iris. The device 242also has an ultrasound transducer integral to the contact lens andcircumferentially positioned with respect to the hollow center of thecontact lens 244. The ring transducer 244 is capable of a largerdetecting region than the needle configuration. Additionally, the ringtransducer 244 is capable of spherical focusing. By example, the ringtransducer can be manufactured from a LiNbO3 single crystal, a compositematerial, PZT and PVDF.

The ring transducer 244 is integral with respect to the contact lens244. It is contemplated that the ring transducer and contact lens can becombined in a variety of physical configurations. The ring transducer244 can be attached to the outer wall 248, attached to the inner wall250, integrated within and completely encompassed by the contact lens,as well as recessed within the contact lens with respect to either theouter wall 248 or inner wall 250. By example, the ring transducer isrecessed and attached to the inner wall 250 of the contact lens 244.When the contact lens 244 is placed upon the eye the transducer does notdirectly touch the eye since it is recessed within the inner wall 250.Similar to the void identified within FIG. 10 a and b, the recess fillswith fluid in the eye, such as tears, and acts as a natural ultrasoniccoupling medium. It is further contemplated that alternativeconfiguration be selected based upon the distance from the retinalregion of interest, the size and shape of the eye, and the specificconfigurations of the transducer selected.

Insulated wires will be used to connect the transducer with an amplifieroutside the contact lens. The amplifier is operatively connected to adigitizer and computer for processing the ultrasonic data.Alternatively, the transmitter can be integrated in the contact lens andthe detected acoustic signal can be transmitted wirelessly.

In another embodiment, the device 242 includes a contact lens and ringtransducer as shown in FIG. 10 a. The device further includes amicroprocessor, memory storage device and wireless transmitter (notshown). The device is placed upon an eye and measures photoacousticwaves generated from an irradiated retinal region of interest. Themicroprocessor digitizes and stores the data within the integratedmemory storage device. After the device 242 is no longer acquiringphotoacoustic data the device is removed from the eye and placed withinclose proximity to a wireless receiver for downloading the photoacousticdata acquired by the device. Alternatively, the wireless transmitter isof sufficient power to transmit the data in real-time to a wirelessreceiver for processing the photoacoustic data and generating functionalinformation and images for the retinal region of interest. It is furthercontemplated that the wireless transmitter is by default nonfunctionalwhile the device 242 is placed upon the eye in order to prevent thetransmission from interfering with data acquisition and health/safetyconcerns.

According to at least one embodiment, the device 242 acquiresphotoacoustic waves generated by an irradiated retina. The photoacousticwaves are amplified and digitized for processing by a computer orsuitable processor that is capable of generating functional ophthalmicimages based at least in part upon the processed photoacoustic waves.

In another embodiment, the transducer is manufactured from a PMN-33%PT/epoxy composite material, which is identified to function within apiezoelectric longitudinal mode. In yet another embodiment, reactive ionetching is utilized to provide crystal post aspect rations and narrowkerfs often needed for high frequency composites with minimal breakage.

It is contemplated that the various systems, methods and embodimentsdescribed herein can be used for the diagnosis and evaluation ofage-related macular degeneration, geography atrophy, diabeticretinopathy, premature retinopathy, glaucoma, ocular tumors, retinaledema, retinal detachment, several types of ischemic retinopathy, braindisorders and Alzheimer's disease.

Before any embodiments of the invention are explained in detail, it isto be understood that the invention is not limited in its application tothe details of construction and the arrangement of components set forthin the following description or illustrated in the following drawings.The invention is capable of other embodiments and of being practiced orof being carried out in various ways.

1. An ophthalmic imaging system comprising: a laser capable ofgenerating a laser beam for irradiating a biological sample; an opticalcoherence tomography (OCT) apparatus capable of generating an OCTprobing light beam; a dual-axis scanner for scanning both the laser beamand OCT probing light beam; an optical apparatus for delivering thelaser beam and OCT probing light beam to the biological sample; and anultrasonic transducer for measuring laser induced ultrasonic waves inthe biological sample.
 2. The system according to claim 1, wherein thelaser system includes a tunable pulse laser.
 3. The system according toclaim 1, and further comprising a processor for generating an imagebased at least in part upon the optical absorption contrast of thebiological sample.
 4. The system according to claim 1, wherein the OCTapparatus includes a broadband light source, a fiber coupler, and aspectrometer for detecting interference signals within a spectraldomain.
 5. The system according to claim 1, wherein the spectral-domainOCT apparatus is fiber-based.
 6. The system according to claim 1, andfurther comprising a photodiode for recording the energy of each laserbeam for use in compensating laser pulse energy instability.
 7. Thesystem according to claim 1, wherein the dual-axis scanner is an x-ygalvanometer scanner or a polygon mirror scanner.
 8. The systemaccording to claim 1, wherein the optical system is capable of producingan optical tuning range from about 300 nm to about 1300 nm.
 9. Thesystem according to claim 4, and further comprising a processor forprocessing imaging data.
 10. The system according to claim 7, whereinthe ultrasonic detector is selected from the group consisting of asingle stationary detector, an array of detectors, and a contact lensintegrated with an ultrasonic detector.
 11. The system according toclaim 9, wherein the processor includes a time based controller capableof synchronizing laser triggering, optical scanning and dataacquisition.
 12. The system according to claim 11, wherein the timebased controller is capable of dynamically registering the timing oflaser triggering and photoacoustic data acquisition, galvanometerscanning and OCT data acquisition.
 13. The system according to claim 12,and further comprising a GUI for displaying a functional and anatomicalimage of the biological sample based at least in part upon the detectedphotoacoustic waves.
 14. The system according to claim 12, and furthercomprising a processor having executable code capable of detecting theboundaries of different anatomical layers.
 15. A method of imaging abiological sample comprising: generating a laser beam capable ofirradiating a region of interest on a biological sample, the laser beamgenerated from a tunable laser system; merging the laser beam with anoptical coherence tomographic probing light to form a merged beam;scanning the merged beam using a dual-axis optical scanner capable oftwo-dimensional raster scanning; recording photoacoustic waves generatedby the irradiated retinal region; recording light reflected off of thebiological sample; controlling the recording and scanning within asingle time base; and generating an ophthalmic image based at least inpart upon scanning the merged beam and recording photoacoustic waves andreflected light.
 16. The method according to claim 15, wherein thebiological sample is selected from the group consisting of an eye,retina and anterior segment of an eye.
 17. The method according to claim15, wherein the image includes retinal vessel morphology.
 18. The methodaccording to claim 15, wherein merging the laser beam with an opticalcoherence tomography probing light is performed by a dichrotic mirror.19. The method according to claim 15, and further comprisingsynchronizing the timing of laser triggering and photoacoustic dataacquisition, optical scanning and OCT data acquisition.
 20. The methodaccording to claim 15, and further comprising three-dimensionallyregistering OCT and photoacoustic images.
 21. The method according toclaim 15, wherein generating the ophthalmic image includes a threedimensional OCT image and photoacoustic vessel morphology.
 22. Themethod according to claim 17, and further comprising measuring theselective absorption of incident photons by hemoglobin and generating ansO₂ level in retinal vessels.
 23. The method according to claim 15, andfurther comprising generating an HbR value based at least in part uponprocessing the photoacoustic signals.
 24. The method according to claim21, and further comprising noninvasively measuring the retinal bloodhemoglobin oxygen saturation.
 25. The method according to claim 20, andfurther comprising generating an HbO₂ value based at least in part uponprocessing the photoacoustic signals.
 26. A method of ophthalmic imagingcomprising: irradiating a retinal region of interest with a laser beamgenerated from a tunable laser; scanning the retinal region of interestusing a dual-axis galvanometer capable of two-dimensional rasterscanning; merging the laser beam with an OCT probing light to form amerged beam; recording photoacoustic waves generated by the irradiatedretinal region; controlling the recording and scanning within a singletime base; and generating an ophthalmic image based at least in partupon the recording and scanning steps.
 27. An OCT-guided laser scanningphotoacoustic microscope comprising: a tunable laser capable ofirradiating a biological sample with a laser beam; a dual-axisgalvanometer for raster scanning the laser beam; a fiber-basedspectral-domain optical coherence tomography (OCT) system capable ofgenerating an OCT probing light beam, the OCT system having aspectrometer for detecting interference signals within a spectraldomain; an optical delivery system for merging the laser beam and OCTprobing light beam and delivering a merged beam to the biologicalsample; an ultrasound transducer integrated within a contact lens fordetecting photoacoustic waves generated by the irradiated biologicalsample; and a GUI for displaying a image of the biological sample basedat least in part upon the photoacoustic waves and spectrometer detectedsignals.